Pet/mr scanners for simultaneous pet and mr imaging

ABSTRACT

In a combined system, a magnetic resonance (MR) scanner includes a magnet configured to generate a static magnetic field at least in a MR examination region from which MR data are acquired. Radiation detectors are configured to detect gamma rays generated by positron-electron annihilation events in a positron emission tomography (PET) examination region. The radiation detectors include electron multiplier elements having a direction of electron acceleration arranged substantially parallel or anti-parallel with the static magnetic field. In some embodiments, the magnet is an open magnet having first and second spaced apart magnet pole pieces disposed on opposite sides of a magnetic resonance examination region, and the radiation detectors include first and second arrays of radiation detectors disposed with the first and second spaced apart magnet pole pieces.

CROSS REFERENCE TO RELATED APPLICATIONS

This is a divisional application of U.S. Ser. No. 12/521,907 filed Jul.1, 2009 which is a national application of PCT application no.PCT/IB2008/050046 filed Jan. 8, 2008 which claims the benefit of U.S.provisional application Ser. No. 60/884,486 filed Jan. 11, 2007, all ofwhich are incorporated herein by reference.

DESCRIPTION

The following relates to the imaging, diagnostic, and related arts. Itfinds particular application in simultaneous magnetic resonance (MR)imaging and positron emission tomography (PET) imaging, and is describedwith particular reference thereto. However, the following finds moregeneral application in acquisition of PET and MR data from a commonsubject performed simultaneously, sequentially, in a time-interleavedfashion, or by some combination thereof, and to diagnostic processesusing same such as imaging, magnetic resonance spectroscopy, and soforth.

MR and PET imaging are imaging modalities that can sometimes providemore information operating in concert than is provided by eithermodality operating alone. To maximize the synergy of combined MR and PETimaging, it would be useful to perform simultaneous MR and PET imaging,or at least to perform MR and PET imaging together over a relativelyshort time interval, for example while the subject remains stationary ona common patient bed with respect to the short time interval. Suchintegrated data acquisition would simplify spatial and optional temporalregistration of images acquired by MR and PET, and would reduce thelikelihood of occurrence of an operatively significant change in thepatient or other subject between the MR and PET imaging dataacquisitions. Other advantages of combined PET/MR imaging include theability to use MR to construct an attenuation map for use in PETimaging, and the use of MR and PET together for motion compensation.

However, construction of a combined PET/MR scanner (sometimes called ahybrid PET/MR scanner) has heretofore been hindered by detrimentaleffect of the magnetic field of the MR acquisition sub-system on thephotomultiplier tube (PMT) detectors of the PET acquisition sub-system.PMT detectors operate by avalanche multiplication of electrons. In atypical PMT arrangement, a photocathode is biased negatively. A photonstriking the photocathode generates an initial burst of one or moreelectrons that travel through vacuum to a first dynode, where theyinduce generation of a larger number of electrons that travel throughvacuum to a second dynode, and so forth, until the avalanche-multipliedelectron burst reaches the anode. For PET applications, the PMT detectoris typically arranged to view a scintillator that generates a burst oflight (i.e., ultraviolet, visible, and/or infrared photons) responsiveto interaction with a 511 keV gamma photon generated by apositron-electron annihilation event.

As the PMT operation is based on travel of electrons (which are chargedparticles) through a vacuum, a force proportional to the cross-productof the electron charge times the electron velocity and the magneticfield is exerted. This force can be represented, for example, in a formsuch as F=qv×B where q is the electron charge, v is the electronvelocity vector, B is the magnetic field vector, and F is the forceexerted on the electron traveling at velocity v by the magnetic field B.The effect of magnetic field on the PMT operation is suitably explainedas follows. The process of electron multiplication is done via anelectrostatic field E in the PMT. If there is also a static magneticfield B present, the force on the electron is given by F=q(E+v×B) and sothe electron no longer accelerates along the direction of theaccelerating electric field E for |v|>0. In consequence the accelerationof the electron and thus the multiplication is disturbed by the presenceof the magnetic field B. For example, if the velocity of the electron iscalculated as a function of time assuming a magnetic field orthogonal tothe accelerating electric field, the velocity of the electron is seen tobe zero at specific points in time, which corresponds to a restart ofthe acceleration. The extreme sensitivity of typical PMT detectors issuch that even the earth's magnetic field (typically about 3×10⁻⁵ T to6×10⁻⁵ T) is sufficient to degrade PMT operation. The small magneticfield of the earth can be compensated by suitable calibration of the PMTdetector. In contrast, a typical MR scanner generates a static (B₀)magnetic field of 0.2 T to 7 T, depending upon the strength of the mainmagnet, with higher-field MR scanners in development. The effect of thismuch larger B₀ magnetic field on the PMT operation cannot be adequatelycompensated by calibration.

One approach for overcoming this problem is to use a solid stateradiation detector that is less sensitive to the magnetic field comparedwith a PMT-based detector. For example, International Publication WO2006/111869 discloses PET/MR scanners that employ solid state siliconphotomultipliers (SiPM) or avalanche photodiodes (APD) as radiationdetectors. This solution involves use of SiPM, APD, or othernon-conventional detectors.

Another approach for overcoming the PMT/MR magnetic field interaction isto move the PMT detectors out of the MR acquisition system, and henceaway from the strong magnetic field. The scintillators remain in or nearthe MR acquisition system, and are coupled with the remote PMT detectorsusing fiber optical connections. This approach retains the conventionalPMT-based detectors, but has the disadvantages of increased systemcomplexity. An additional disadvantage for time-of-flight (TOF) PETimaging is that the fiber optical connections add significant light lossand transit time spread that adversely affect TOF calculations.Wavelength-dependent dispersion further complicates these TOFcalculations and limits the temporal resolution achievable using fiberoptical connections.

The following provides new and improved apparatuses and methods whichovercome the above-referenced problems and others.

In accordance with one aspect, a combined magnetic resonance (MR) andpositron emission tomography (PET) data acquisition system is disclosed.A magnetic resonance scanner includes a magnet configured to generate astatic magnetic field at least in a magnetic resonance examinationregion. The magnetic resonance scanner is configured to acquire magneticresonance data from the magnetic resonance examination region. Radiationdetectors are configured to detect gamma rays generated bypositron-electron annihilation events in a positron emission tomographyexamination region. The radiation detectors include electron multiplierelements having a direction of electron acceleration arrangedsubstantially parallel or anti-parallel with the static magnetic field.

In accordance with another aspect, a data acquisition method isdisclosed. Magnetic resonance data are acquired using the magneticresonance scanner of the combined MR and PET data acquisition system ofthe preceding paragraph. PET data are acquired using the radiationdetectors of the combined MR and PET data acquisition system of thepreceding paragraph. The MR and PET acquiring operations are performedsimultaneously.

In accordance with another aspect, a combined magnetic resonance (MR)and positron emission tomography (PET) data acquisition system isdisclosed. An open magnetic resonance scanner has first and secondspaced apart magnet pole pieces disposed on opposite sides of a magneticresonance examination region and is configured to generate a staticmagnetic field at least in the magnetic resonance examination regionhaving a magnetic field direction directed from the first magnet polepiece to the second magnet pole piece. First and second arrays ofradiation detectors are disposed with the first and second spaced apartmagnet pole pieces and are configured to detect gamma rays generated bypositron electron annihilation events in a positron emission tomographyexamination region that at least partially overlaps the magneticresonance examination region.

In accordance with another aspect, a data acquisition method isdisclosed. Magnetic resonance data are acquired using the open magneticresonance scanner of the combined MR and PET data acquisition system ofthe preceding paragraph. Simultaneously with the acquiring of magneticresonance data, PET data are acquired using the first and second arraysof radiation detectors of the combined MR and PET data acquisitionsystem of the preceding paragraph.

In accordance with another aspect, a data acquisition method isdisclosed. A static magnetic field is generated at least within amagnetic resonance (MR) examination region. MR data are acquired fromthe MR examination region. Positron emission tomography (PET) data areacquired from a PET examination region at least partially overlappingthe MR examination region. The acquiring of PET data includes amplifyingradiation detection event signals by electron multiplication processesconfigured such that cross products of velocities of acceleratingelectrons of the electron multiplication processes and the staticmagnetic field are substantially nulled.

In accordance with another aspect, a positron emission tomography (PET)data acquisition system is disclosed. An array of scintillators surroundan examination region and are configured to generate scintillationevents responsive to interaction with gamma rays. A plurality ofmicrochannel photomultipliers are optically coupled with thescintillators to detect the scintillation events. A processor isconfigured to reconstruct projection data or spatially localizedprojection data derived from the detected scintillation events into areconstructed positron emission tomographic image.

In accordance with another aspect, a combined magnetic resonance (MR)and positron emission tomography (PET) data acquisition system isdisclosed, including: an open magnetic resonance scanner including firstand second spaced apart magnet pole pieces, the open magnetic resonancescanner configured to acquire a magnetic resonance image of a subject; adiscontinuous array of radiation detectors incompletely encircling saidsubject and configured to detect 511 keV gamma rays emanating from saidsubject; and electronics configured to (i) acquire time of flightlocalized positron emission tomography projection data using thediscontinuous array of radiation detectors and (ii) reconstruct apositron emission tomography image of the subject from the time offlight localized positron emission tomography projection data.

In accordance with another aspect, an imaging method is disclosed,comprising: acquiring time of flight localized positron emissiontomographic projection data from a subject using an array of radiationdetectors that incompletely encircle the subject; and reconstructing animage of the subject from the acquired time of flight localized positronemission tomographic projection data, wherein information missing due tothe incomplete encirclement of the subject by the array of radiationdetectors is compensated by additional information provided by the timeof flight localization.

In accordance with another aspect, a radio frequency coil is disclosedfor use in magnetic resonance imaging, the radio frequency coilcomprising: a resonant structure having a resonance frequency consonantwith a magnetic resonance frequency; a plurality of scintillatorssecured with the resonant structure; and optical detectors arranged todetect scintillation events emanating from the scintillators.

One advantage resides in facilitating concurrent PET and MR dataacquisition.

Another advantage resides in providing a PET/MR scanner employingphotomultiplier tube (PMT)-based radiation detectors.

Another advantage resides in providing an open PET/MR scanner.

Another advantage resides in facilitating integration of PET detectorsincluding photomultiplier tubes into an MR scanner.

Still further advantages of the present invention will be appreciated tothose of ordinary skill in the art upon reading and understand thefollowing detailed description.

FIG. 1 diagrammatically shows a combined magnetic resonance (MR) andpositron emission tomography (PET) data acquisition system including anopen MR scanner portion.

FIG. 2 diagrammatically shows illustrative image processing and displaycomponents coupled with the hybrid data acquisition system of FIG. 1.

FIG. 3 diagrammatically shows a portion of the lower planar array ofradiation detectors of the PET data acquisition portion of the system ofFIG. 1.

FIG. 4 diagrammatically shows an embodiment of the electron multiplierelements of the lower planar array of radiation detectors of FIG. 4, inwhich the electron multiplier elements include microchannel platephotomultipliers.

FIG. 5 diagrammatically shows a side view of a planar array of radiationdetectors employing a block readout configuration.

FIGS. 6 and 7 show portions of alternative embodiments of the lowerplanar array of radiation detectors of the PET data acquisition portionof the system of FIG. 1, in which the scintillators are wholly (FIG. 6)or partially (FIG. 7) spatially integrated with the radio frequencycoil.

FIGS. 8 and 9 diagrammatically show perspective and end sectional views,respectively, of a combined magnetic resonance (MR) and positronemission tomography (PET) data acquisition system including a closedbore-type MR scanner portion.

FIG. 10 diagrammatically shows a portion of the array of radiationdetectors of the PET data acquisition portion of the system of FIGS. 8and 9 viewed from the bore outward.

With reference to FIG. 1, a combined or hybrid magnetic resonance (MR)and positron emission tomography (PET) data acquisition system 8 is setforth as an illustrative example. A magnetic resonance scanner includesan open magnet 10 including a plurality of conductor coil windings 11(diagrammatically depicted in FIG. 1 by boxes with crossing lines) thatare energized by electrical current to generate a static magnetic fieldB₀ at least within a magnetic resonance examination region 12. Themagnetic resonance examination region 12 is indicated by a dashedborderline, and is circular in the illustrated embodiment; however, themagnetic resonance examination region may in general be circular,elliptical, or otherwise-shaped. The conductor coil windings 11 may besuperconducting or resistive windings; in the former case the windingsare typically disposed in a cryogenic container or other cooling system(not shown). The illustrated magnet 10 is an open magnet having an firstand second pole pieces, such as the illustrated upper pole piece 14 anda lower pole piece 15, separated by a gap including the magneticresonance examination region 12. The illustrated open magnet 10 is avertical magnet producing the static magnetic field B₀ having a vertical(e.g., up or down) magnetic field direction. The illustrated staticmagnetic field B₀ is directed from top-to-bottom, so that the upper polepiece 14 is a north pole while the lower pole piece 15 is a south pole.The opposite polarity is also suitable. In other embodiments, the openmagnet may be otherwise oriented to produce a horizontal orotherwise-oriented static magnetic field. In the embodiment illustratedin FIG. 1, at least the upper pole piece 14, and typically also thelower pole piece 15, are supported and spaced apart by vertical framepieces 16, 17. The frame pieces 16, 17 are illustrative examples; othermechanical support arrangements are conceivable.

The magnetic resonance scanner also includes a magnetic field gradientassembly, which in the illustrative embodiment of FIG. 1 includes upperand lower gradient coil windings 20, 21 that cooperatively superimposemagnetic field gradients on the static B₀ magnetic field responsive toselective energizing of selected gradient coil windings. Optionally, themagnetic field gradient coil, magnet, or both may include other featuresnot shown for forming, stabilizing, and dynamically adjusting themagnetic field, such as passive ferromagnetic shims, active shimmingcoils, or so forth. The magnetic resonance scanner further includes aradio frequency excitation and reception system, such as an illustratedbuilt-in radio frequency coil (indicated diagrammatically) having upperand lower generally planar portions 24, 25 and including a radiofrequency screen or shield including upper and lower screen or shieldportions 26, 27 (indicated diagrammatically by dashed lines). The radiofrequency system includes at least one component, such as theillustrated radio frequency coil 24, 25, that can be energized at asuitable radio frequency to excite magnetic resonance in a subjectdisposed in the magnetic resonance examination region 12, and typicallyincludes at least one component, such as the illustrated radio frequencycoil 24, 25, that can operate as a radio frequency receiver to receiveor detect magnetic resonance emanating from the magnetic resonanceexamination region 12. In some embodiments, different coils are used forthe excitation and reception operations. For example, the built-in coil24, 25 may be used to excite magnetic resonance, and a different, localcoil (not shown) may be positioned over or close to the subject in themagnetic resonance examination region 12 to detect magnetic resonance.It is contemplated for the same magnetic resonance scanner to beconfigurable in different ways using different combinations of built-incoils, local coils, or both.

With continuing reference to FIG. 1 and with brief reference to FIG. 2,in a magnetic resonance imaging session example, the radio frequencyexcitation system 24, 25 is used to excite magnetic resonance in thesubject, while magnetic field gradients applied by the gradient system20, 21 before, during, or after the magnetic resonance excitationspatially localize the excited magnetic resonance or encode the excitedmagnetic resonance by frequency-encoding, phase-encoding, or so forth.The excited and spatially encoded magnetic resonance is received usingthe radio frequency reception system 24, 25 (in this example, the samecoil 24, 25 is used for both excitation and reception, but in generaldifferent coils may be used) and received magnetic resonance samples arestored in a magnetic resonance sampling storage 30. A magnetic resonancereconstruction processor 32 applies a suitable reconstruction algorithmto reconstruct the magnetic resonance samples to form a reconstructedimage that is stored in a magnetic resonance images memory 34. Thereconstruction processor 32 applies a reconstruction algorithm thatcomports with the selected spatial encoding used in generating themagnetic resonance data. For example, a Fourier transform reconstructionalgorithm may be suitable for reconstructing Cartesian-encoded magneticresonance data.

With continuing reference to FIGS. 1 and 2, the combined or hybrid MRand PET data acquisition system 8 further includes radiation detectorsfor performing PET data acquisition. In the illustrative example of FIG.1, the radiation detectors include first and second generally planararrays 40, 41 of radiation detectors. As will be described, each of theillustrated generally planar detector arrays 40, 41 includes ascintillator layer 56 and a layer of electron multiplier-based photondetectors 60; however, other detector configurations are contemplatedfor use as the arrays 40, 41. The first generally planar array 40 ofradiation detectors is arranged between the first magnet pole piece 14and the first portion 24 of the radio frequency coil. The secondgenerally planar array 41 of radiation detectors is arranged between thesecond magnet pole piece 15 and the second portion 25 of the radiofrequency coil. Each of the generally planar arrays 40, 41 of radiationdetectors is configured to detect 511 keV gamma rays that are emitted bypositron-electron annihilation events. As is known in the art, when anelectron and positron meet, they annihilate, emitting two 511 keV gammarays that are oppositely directed in accord with the principle ofconservation of momentum. (There may be a slight deviation from preciseopposition due to carried impulse of the electron and positron prior toannihilation, again in accordance with conservation of momentum). In PETdata acquisition, two substantially simultaneous 511 keV gamma raydetection events are presumed to have originated from the samepositron-electron annihilation event, which is therefore locatedsomewhere along the “line of response” connecting the two substantiallysimultaneous 511 keV gamma ray detection events. This line of responseis also sometimes called a projection, and the collected PET data isreferred to as projection data.

In conventional PET, substantially simultaneous 511 keV gamma raydetection events are defined as two 511 keV gamma ray detection eventsoccurring within a selected short time window, such as within onenanosecond of each other. Due to the variable annihilation position withrespect to the detector elements a small (e.g., sub-nanosecond) timedifference between the substantially simultaneous gamma photon detectionevents occurs. A related technique, called time-of-flight PET orTOF-PET, takes advantage of this small time difference to furtherlocalize the positron-electron annihilation event along theline-of-response. In general, the annihilation event occurred along theprojection at a point closer to the gamma ray detection event thatoccurred first. If the two gamma ray detection events occursimultaneously within the time resolution of the detectors, then theannihilation event occurred at the midpoint of the projection. Theachievable TOF spatial localization along the projection is dependentupon the time resolution, jitter, and other temporal characteristics ofthe radiation detectors. A projection with TOF spatial localizationalong the projection is referred to herein as a spatially localizedprojection.

With continuing reference to FIGS. 1 and 2, the radiation detectorarrays 40, 41 of the hybrid system 8 are used to acquire PET or TOF-PETdata (that is, projection data or spatially localized projection data).The gamma ray detection events are processed by a PET digitization unit42 that performs time-to-digital conversion (TDC) and analog-to-digitalconversion (ADC) of detection events, and a singles processing unit 43that performs clustering, energy estimation, timestamping, andpositioning. The singles processing unit 43 optionally filters outdetections that are outside of a selection energy window centered abouton the expected 511 keV gamma ray energy. In some embodiments, theradiation detectors are pixelated, so that the spatial localization ofthe gamma ray detection events defining the projection correspond to apixel size (i.e., physical size) of the radiation detectors of theradiation detector arrays 40, 41. In other embodiments, clustering isapplied by a block readout algorithm such as Anger logic or the like toprovide further spatial localization refinement of the gamma raydetection events defining the projection. A coincidence detectionprocessor 44 employs temporal windowing to identify gamma ray detectionevents that occurred substantially simultaneously, and hence likelycorrespond to a common positron-electron annihilation event and hencedefine a projection or line of response.

For TOF processing, the singles processing 43 and coincidence detectionprocessing 44 can be swapped or interleaved so that the time differencebetween the identified substantially simultaneous or coincidentdetection events can be used to spatially localize the positron-electronannihilation event along the projection or line of response. Thelocalization optionally takes the form of a Gaussian or otherstatistical localization. In one approach, clustered events are input toa three-dimensional position estimation algorithm (for example,implemented using look-up tables) to give local spatial coordinates ofthe corresponding block detector. Centralized coincidence processingidentifies the lines of response. Event information is optionallycorrelated from coincident radiation detection events to provide refinedpositioning, and is optionally correlated from neighboring blocks forevent synthesis and randoms reduction using knowledge of the Comptonscatter cone.

The resulting PET or TOF-PET data are stored in a PET data storage 45. APET reconstruction processor 47 processes the projection or localizedprojection data using a suitable reconstruction algorithm to generate areconstructed image that is stored in a PET images storage 48. Forexample, a filtered backprojection algorithm or iterative reconstructionalgorithm can be employed. The system of FIG. 1 includes an openmagnetic resonance scanner and a discontinuous radiation detector array40, 41 for PET that does not completely encircle the subject. Theincomplete encirclement can lead to imaging artifacts due to “missing”projections or lines of response. For example, in the system of FIG. 1no perfectly horizontal projections are collected, and so informationordinarily provided by such horizontal projections about verticalposition is unavailable. Advantageously, if time-of-flight PET data areacquired and reconstructed then the time-of-flight localization providesadditional information that compensates for the information that is lostby incomplete encirclement. As a qualitative example, the aforementionedmissing information about vertical position can be compensated by aTOF-localized vertical projection, since the temporal localization ofthe electron-positron annihilation event along the vertical projectionprovides information about its vertical position.

The MR and PET acquisitions are optionally performed concurrently. Aswill be discussed, this concurrent acquisition is enabled by configuringthe hybrid acquisition system 8 to reduce or eliminate detrimentalinteraction between the radiation detector arrays 40, 41 of the PETacquisition sub-system and the static magnetic field B₀ generated by themagnet 10 of the magnetic resonance acquisition sub-system.Alternatively or additionally, MR and PET acquisition can be donesequentially (e.g., MR first followed by PET, or vice versa) or can beinterleaved (e.g., acquire an MR image, a PET image, then translate thepatient using a motorized couch or other support to a next imagingposition, and repeat the MR and PET acquisitions). An imagesregistration processor 50 optionally spatially registers and optionallytemporally registers the reconstructed MR and PET images, and the imagesas so registered are suitably displayed on an images display device 52,or stored, rendered using suitable two- or three-dimensional renderingsoftware, or otherwise processed.

With reference to FIG. 1 and with further reference to FIG. 3, anembodiment of the radiation detector arrays 40, 41 of the PETacquisition sub-system is described. FIG. 3 depicts a small portion ofthe lower (or second) radiation detector array 41. The first and secondgenerally planar arrays 40, 41 of radiation detectors are symmetrical inthe illustrated embodiment, and each contain a planar array ofscintillators 56 and a planar array of electron multiplier elements 60.For time-of-flight PET data acquisition, the scintillators 56 aresuitably crystals of LaBr₃, LYSO, LGSO, CeBr₃, LuI, LSO, or so forth.The scintillators 56 are advantageously selected to be relativelytransparent to the magnetic resonance acquisition, and should providefast scintillation decay times for time-of-flight measurements. Otherscintillator materials can also be used. The electron multiplierelements 60 are optically coupled with the scintillators 56 and producean amplified electrical signal responsive to a burst of light in theoptically coupled scintillator through an electron multiplicationprocess. Although not shown, in some embodiments a planar light guide isdisposed between the scintillators 56 and the electron multipliers 60.Although the shield 27 is illustrated at the top of the scintillators,it is to be appreciated that it can be positioned at the bottom end orin between the top and bottom.

With continuing reference to FIG. 3 and with further reference to FIG.4, in the illustrated embodiment the electron multiplier elements 60include microchannel plate photomultipliers. Commercially availablemicrochannel plate photomultipliers are sometimes referred to as“microchannel plate photomultiplier tubes” or “MCP-PMT's”. As best seenin FIG. 4, these devices have a microchannel plate 62 of a materialhaving high photoelectric and secondary election emissioncharacteristics. A plurality of parallel microchannels 64 are formedinto the microchannel plate 62. The microchannels 64 are pores, tubes,or other parallel narrow openings. Electrodes 66, 67 generate anelectrical bias across the length of the microchannels 64. An electrongenerated near the open end of a microchannel, for example byphotoelectric effect responsive to impingement of a photon on aphotocathode (not shown), is driven down the microchannel 64 due to theelectrical bias on the electrodes 66, 67. As it travels through themicrochannel 64, the electron occasionally collides with sides of themicrochannel 64. Secondary electron emissions are produced by suchcollisions, thus providing electron multiplication. At the bottom of themicrochannel, the multiplied electrons form an amplified electron signalthat is collected by an anode (not shown). Typically, the electronmultiplier element 60 together with a suitable photocathode and anodeare housed in a vacuum-tight housing. In some embodiments, the materialof the electron multiplier 60 may provide the material for thephotoelectric effect to generate the initial electrons. In someembodiments the microchannels may be tilted, zigzagged (as shown in FIG.4), curved, or otherwise-configured to enhance the frequency ofelectron-microchannel wall collisions. It is also contemplated to useother biasing arrangements than the illustrated arrangement.

Microchannel plate photomultipliers are typically used in photography,night vision goggles, or other applications in which the naturaltwo-dimensional configuration of the microchannels readily provides anaddressable two-dimensional optical detector array with built-inamplification. The radiation detector arrays 40, 41 make use of adifferent advantage of the microchannel plate arrangement. Themicrochannel plates 62 of the microchannel plate photomultipliers 60 arearranged with the microchannels 64 parallel with the static magneticfield B₀. In this configuration the microchannels 64 are arranged toeffectively view the scintillators 56 directly and, additionally, adirection of electron acceleration a_(e) along the axes of themicrochannels 64 is arranged parallel or anti-parallel with the staticmagnetic field B₀. With this geometry, the cross products of thevelocities of accelerating electrons of the electron multiplicationprocesses in the microchannels 62 and the static magnetic field B₀ aresubstantially zero. As a result, the effect of the static magnetic fieldB₀ on the microchannel plate photomultipliers 60 is substantiallyreduced or eliminated, even though the microchannel platephotomultipliers 60 are disposed within the strong static magnetic fieldB₀.

With reference to FIG. 5, in some embodiments a block readoutconfiguration is used, in which each scintillator 56 is a block that isviewed by a microchannel plate photomultiplier or an array of SiPMs. InFIG. 5, microchannel plate photomultipliers 60 are used, with themicrochannels 64 diagrammatically indicated. By applying a suitableblock readout algorithm as part of the singles processing 43,substantial spatial resolution is achievable. For example, withavalanche photo diodes (APDs), a spatial resolution of about 2 mm hasbeen shown to be achievable within a 20×10×10 mm³ LSO scintillator blockusing block readout processing. See Maas et al., ExperimentalCharacterization of Monolithic-Crystal Small Animal PET Detectors ReadOut by APD Arrays, IEEE Trans. Nucl. Sci. vol. 53 no. 3, pp. 1071-77(2006) which is incorporated herein by reference in its entirety. In theillustrative arrangement of FIG. 5, a first printed circuit board 68supports electronics, such as microprocessors, application-specificintegrated circuits (ASIC's), memory chips, or so forth, that embody theevent processors 42, 43 of FIG. 1, or some portion thereof. For example,in some embodiments the electronics of the first printed circuit board68 embody the TDC/ADC units 42, while the singles processing 43 togetherwith the coincidence detection 44 are performed remotely. In thearrangement illustrated in FIG. 5, a second printed circuit board 69defines a readout backplane supporting suitable electronics and printedcircuitry to read and route off-board the signals generated by theelectronics of the first printed circuit board 68.

With reference to FIGS. 1 and 3, the upper and lower photodetectorarrays 40, 41 are arranged to view a positron emission tomography (PET)examination region 70 through respective upper and lower generallyplanar portions 24, 25 of the radio frequency coil. The PET examinationregion 70 is indicated by a dotted borderline, and is circular in theillustrated embodiment; however, the PET examination region may ingeneral be circular, elliptical, or otherwise-shaped. In theillustrative example of FIG. 3, the radio frequency coil includes aprinted circuit board 72 having printed conductors, such as theillustrated printed conductor portion 74, formed on a front surface. Theradio frequency shield or screen 27 can be a free-standing wire mesh, ora conductor mesh printed or otherwise disposed on a printed circuitboard 72 or other substrate, or so forth. The upper planar portion 24 ofthe radio frequency coil suitably has the same configuration. In thearrangement shown, the radio frequency coil 24, 25 should besubstantially transparent to 511 keV gamma rays, so that the radiofrequency coil 24, 25 does not block the view of the PET examinationregion 70 by the photodetector arrays 40, 41. In the illustratedembodiment, the PET examination region 70 is slightly smaller than theMR examination region 12. In other embodiments, the MR examinationregion may be the smaller region, or the centroid of the MR examinationand the centroid of the PET examination region may be relatively offsetfrom one another, or so forth. If there is at least some overlap of theMR and PET examination regions, then acquisition of both MR and PETimaging data of the same portion of the subject is possible withoutmoving the subject between the MR and PET acquisitions. Indeed, it iscontemplated to acquire PET and MR imaging data from the same region ofthe subject simultaneously, or in an interleaved fashion.

With reference to FIG. 1, the photodetector arrays 40, 41 furtherinclude respective suitable circuitry 80, 81 for reading the electronmultiplier elements 60. For example, the circuitry 80, 81 may includethe PET digitization processing 42 and optionally the singles processing43 for digitizing the energy of the amplified signals indicative ofgamma ray detection events. The event information can either beself-triggered or common triggered by using a locally installed neighborlogic or analog circuitry. The processing units collecting this dataoptionally perform offset, gain, time-walk, or other corrections, aswell as optional sorting and optional event clustering. In theillustrated open MR scanner configuration, passthroughs in the magnet 10provide conduits 82, 83 for conveying electrical power, optionaldetector cooling, signal transmission, or so forth. In some embodiments,some of the circuitry 80, 81 may be located remote from the electronmultiplier elements 60.

An advantage of the configuration shown in FIGS. 1 and 3 iscompactness—the planar photodetector arrays 40, 41 do not take up a lotof space in the gap between the north and south magnet poles 14, 15.

With reference to FIG. 6, a modified embodiment is illustrated in whichthe planar photodetector arrays 40, 41 take up still less space in thegap between the north and south magnet poles 14, 15. In this embodiment,the scintillators 56 are integrated into a modified radio frequency coil25′. For example, the scintillators 56 may be incorporated into theprinted circuit board 72 that supports the conductors 74 of the radiofrequency coil.

In some embodiments, the scintillators 56 are formed as pixels sized toalign with gaps of a conductive mesh optionally forming the radiofrequency shield or screen 27. In some embodiments, the wire meshincorporates resonant circuits that provide a high impedance forgradient fields produced by the gradient coils 21, and low impedance forthe radio frequency field generated by the radio frequency coil. For astatic magnetic field B₀ greater than or about 1.5 Tesla, the diameterof wires of a conductive mesh forming the radio frequency shield orscreen can be less than 0.1 millimeters. If a planar light guide (notshown) is inserted between the scintillators 56 and the electronmultiplier elements 60, then the light guide can be used as a substratefor supporting the wire mesh and the wire mesh can define alignmentfeatures for positioning blocks of the scintillators 56.

With reference to FIG. 7, an intermediate embodiment is illustrated inwhich the planar photodetector arrays 40, 41 take up an intermediateamount of space in the gap between the north and south magnet poles 14,15. In this embodiment, the scintillators 56 are partially integratedinto a modified radio frequency coil 25″. For example, the scintillators56 may be protrude into a recess of the printed circuit board 72 thatsupports the conductors 74 of the radio frequency coil.

With reference to FIGS. 8 and 9, a hybrid or combined magnetic resonance(MR) and positron emission tomography (PET) data acquisition system 108is shown. A magnetic resonance scanner includes a closed bore-typescanner housing 109 that contains a solenoidal-type magnet 110 having aplurality of conductor coil windings arranged to generate a staticmagnetic field B₀ within and parallel with the axis of a generallycylindrical bore opening 112 defined by the housing. The magnet may be asuperconducting or resistive magnet; in the former case the magnet istypically disposed in a cryogenic container or other cooling systemwithin the bore-type scanner housing 110. The scanner housing 110 alsocontains a magnetic field gradient assembly 120, such as magnetic fieldgradient coils that cooperatively superimpose magnetic field gradientson the static B₀ magnetic field responsive to selective energizing ofselected gradient coil windings. Optionally, the magnetic field gradientcoil, magnet, or both may include other features not shown for forming,stabilizing, and dynamically adjusting the magnetic field, such aspassive ferromagnetic shims, active shimming coils, or so forth.

The magnetic resonance scanner further includes a radio frequencyexcitation and reception system, such as an illustrated built-inbirdcage or transverse electromagnetic (TEM)-type “whole-body” radiofrequency coil having rungs or rods 124 and a radio frequency screen orshield 126. In a birdcage-type coil, the rungs 124 are connected byend-rings 125 (one end-ring 125 is depicted as an example in FIG. 8; noend-rings are shown in FIG. 9) to generate and optionally detectmagnetic resonance over a substantial region of the bore opening 112. Ina TEM-type whole-body coil, the end-rings 125 are omitted, and the rods124 are connected at their ends to the radio frequency shield or screen126 to generate and optionally detect magnetic resonance of asubstantial region of the bore opening 112. In either a birdcage orTEM-type coil, the radio frequency shield or screen 126 is typically athin cylindrical conductive mesh or a resonant screen. In otherembodiments, one or more local radio frequency coils (not shown) usedexclusively or in conjunction with the whole-body coil 124, 126 formagnetic resonance excitation and/or detection.

With continuing reference to FIGS. 8 and 9, the combined or hybrid MRand PET data acquisition system 108 further includes radiation detectors140 for performing PET data acquisition. In the illustrative example ofFIGS. 8 and 9, these radiation detectors 140 are arranged as ashort-axis cylindrical array centered around the isocenter of the MRmagnet. In FIG. 8, the radiation detectors 140 are shown in phantom forillustrative purposes, but it is to be understood that the radiationdetectors 140 are typically blocked from view by the scanner housing109. The cylindrical array of radiation detectors 140 surround and arecoaxial with the bore opening 112 of the MR scanner. As shown in FIG. 9,in the illustrated embodiment the radiation detectors 140 are at aradius (i) smaller than the radius of the gradient assembly 120 and (ii)larger than the radius of the radio frequency screen or shield 126. Thatis, the radiation detectors 140 are placed between the gradient assembly120 and the radio frequency screen or shield 126. Alternatively, theradiation detectors 140 could be arranged inside of (i.e., at a smallerradius than) the shield 126 and outside of (i.e., at a larger radiusthan) the rungs or rods 124. In other embodiments, the radiationdetectors 140 are arranged inside of the shield 126 and inside of therungs or rods 124. In other embodiments, the radiation detectors 140 arearranged at about the same radius as the rungs or rods 126, for examplebetween the rungs or rods. However they are arranged, the radiationdetectors 140 should have a substantially unimpeded view of theexamination region within the bore opening 112 respective to 511 keVgamma rays; intervening components should be either narrow (e.g., therungs or rods 124) or substantially transparent to 511 keV gamma rays(e.g., the radio frequency screen or shield 126).

With continuing reference to FIGS. 8 and 9 and with further reference toFIG. 10, an embodiment of the radiation detectors 140 of the PETacquisition sub-system is described. FIG. 10 depicts a small portion ofthe array of radiation detectors 140 including ten scintillator blocks156 and corresponding electron multiplier elements 160. The illustratedelectron multiplier elements 160 include microchannel platephotomultipliers, such as commercially available microchannel platephotomultipliers sometimes referred to as “microchannel platephotomultiplier tubes” or “MCP-PMT's”. The electron multiplier elements160 are suitably similar to the electron multipliers 60 depicted in FIG.4, and include microchannels 164. As in the case of the acquisitionsystem 8, the microchannel plate photomultipliers 160 of the acquisitionsystem 108 are preferably arranged with their direction of electronacceleration a_(e) parallel or anti-parallel with the static magneticfield B₀. For the configuration of radiation detectors 140 illustratedin FIG. 10, the top layer the direction of electron acceleration a_(e)is anti-parallel with the direction of the static magnetic field B₀,while for the bottom layer the direction of electron acceleration a_(e)is parallel with the direction of the static magnetic field B₀. In FIG.10, the microchannels 164 are diagrammatically indicated and arearranged parallel with the static magnetic field B₀ so that the electricfield generated by the biasing of the electrodes of the microchannelplate photomultipliers 160 along the length of the microchannels 164produces the electron acceleration a_(e) oriented parallel oranti-parallel with the static magnetic field B₀. The resulting geometricarrangement has the microchannel plate photomultipliers 160 arranged atthe sides of their respective scintillators 156. This side-by-sidearrangement results in gaps between the scintillators 156 to accommodatethese microchannel plate photomultipliers 160 and associated electronicsor circuitry. To address these gaps, the radiation detectors 140 shownin FIG. 10 are arranged in a double layer with the scintillators 156 inthe second layer offset from those of the first layer to fill in thegaps. This approach is suitable if the microchannel platephotomultipliers 160 are substantially transparent to 511 keV gammarays. More generally, the scintillators 156 and microchannel platephotomultipliers 160 are optionally interleaved in one or more detectorlayers, with each detector layer having a surface normal orientedtransverse to the direction of the static magnetic field B₀. For acylindrical magnet, the one or more interleaved detector layers eachdefine a generally cylindrical shell arranged coaxially with thegenerally cylindrical magnet 110. If the microchannel platephotomultipliers are not sufficiently transparent, then a single layercan be used, albeit with some loss of resolution.

In the illustrated embodiment, each scintillator 156 is a block that isviewed from the side by an 8×8 array of microchannel platephotomultipliers 160, and Anger logic or another suitable block readoutalgorithm is used to enhance spatial resolution based on the relativesignals received by the elements of the 8×8 array of microchannel platephotomultipliers 160. In other embodiments, pixelated scintillators maybe used, or block scintillators may be used with viewing microchannelplate photomultiplier arrays of other sizes. Suitable circuitry 180 forreading the electron multiplier elements 160 is integrated into theradiation detectors 140 as application-specific integrated circuitry(ASIC) or microprocessor components. For example, the circuitry 180 mayinclude analog-to-digital converters (ADC) for digitizing the energy ofthe amplified signals indicative of gamma ray detection events,time-to-digital converters (TDC) for generating digital timestamps forgamma ray detection events, and so forth. The event information caneither be self-triggered or common triggered by using a locallyinstalled neighbor logic or analog circuitry. The processing unitscollecting this data optionally perform offset, gain, time-walk, orother corrections, as well as optional sorting and optional eventclustering. Optionally, the circuitry 180 includes portions or all ofthe PET events processing 42, 43.

Like the acquisition system 8, the acquisition system 108 is optionallyconfigured for simultaneous MR and PET imaging if the MR and PETexamination regions at least partially overlap. Additionally oralternatively, MR and PET imaging may be performed in succession or maybe interleaved, for example on a slice-by-slice basis.

The illustrated embodiments employ electron multiplier elements 56, 156that are advantageously arranged with the direction of electronacceleration a_(e) substantially parallel or anti-parallel with amagnetic field direction of the static magnetic field B₀ generated bythe magnetic resonance scanner. In other contemplated embodiments,detectors that are relatively insensitive to magnetic field aresubstituted for the electron multiplier elements. Such detectors mayinclude, for example, avalanche photodiodes (APD's) or solid statesilicon photomultipliers (SiPM). Some suitable SiPM detectors aredescribed in International Publication WO 2006/111869, and inInternational Publication WO 2006/111883. Optionally, the radiationdetectors 40, 41, 140 (or selected units or portions thereof) aresurrounded radio frequency-tight and light-tight shielding (not shown)to protect the detectors and electronics. Mu-metal shielding (not shown)is optionally included as well. The illustrated open and bore-typemagnetic resonance scanners are illustrative examples, and other typesof MR scanners are contemplated.

1. A data acquisition method comprising: generating a static magneticfield at least within a magnetic resonance (MR) examination region;acquiring MR data from the MR examination region; and acquiring positronemission tomography (PET) data from a PET examination region at leastpartially overlapping the MR examination region, the acquiring of PETdata including amplifying radiation detection event signals by electronmultiplication processes configured such that cross products ofvelocities of accelerating electrons of the electron multiplicationprocesses and the static magnetic field are substantially nulled.
 2. Thedata acquisition method as set forth in claim 1, wherein the acquiringof MR data and the acquiring of PET data are performed concurrently. 3.A positron emission tomography (PET) data acquisition system comprising:an array of scintillators surrounding an examination region andconfigured to generate scintillation events responsive to interactionwith gamma rays; a plurality of microchannel photomultipliers opticallycoupled with the scintillators to detect the scintillation events; and aprocessor configured to reconstruct projection data or spatiallylocalized projection data derived from the detected scintillation eventsinto a reconstructed positron emission tomographic image.
 4. Thepositron emission tomography (PET) data acquisition system as set forthin claim 3, further including: a non-natural magnetic field, themicrochannel photomultipliers arranged with their microchannels orientedparallel with the non-natural magnetic field.
 5. The positron emissiontomography (PET) data acquisition system as set forth in claim 4,further including: a magnetic resonance (MR) scanner including a magnetgenerating the non-natural magnetic field.
 6. The positron emissiontomography (PET) data acquisition system as set forth in claim 5,wherein the array of scintillators and the microchannel photomultipliersare mounted on the MR scanner to define a hybrid PET/MR scanner whichboth receive data from at least partially overlapping examinationregions.
 7. A combined magnetic resonance (MR) and positron emissiontomography (PET) data acquisition system comprising: an open magneticresonance scanner including first and second spaced-apart magnet polepieces, the open magnetic resonance scanner configured to acquire amagnetic resonance image of a subject; a discontinuous array ofradiation detectors incompletely encircling said subject and configuredto detect 511 keV gamma rays emanating from said subject; andelectronics configured to (i) acquire time-of-flight localized positronemission tomography projection data using the discontinuous array ofradiation detectors and (ii) reconstruct a positron emission tomographyimage of the subject from the time-of-flight localized positron emissiontomography projection data.
 8. An imaging method comprising: acquiringtime-of-flight localized positron emission tomographic projection datafrom a subject using an array of radiation detectors that incompletelyencircle the subject; and reconstructing an image of the subject fromthe acquired time-of-flight localized positron emission tomographicprojection data, wherein information missing due to the incompleteencirclement of the subject by the array of radiation detectors iscompensated by additional information provided by the time-of-flightlocalization.
 9. The imaging method as set forth in claim 8, wherein theprojection data are acquired intermittently, the imaging method furtherincluding: acquiring magnetic resonance data during time gaps betweenthe intermittent projection data acquisitions; and reconstructing animage of the subject from the acquired magnetic resonance data.
 10. Aradio frequency coil for use in magnetic resonance imaging, the radiofrequency coil comprising: a resonant structure having a resonancefrequency consonant with a magnetic resonance frequency; a plurality ofscintillators secured with the resonant structure; and optical detectorsarranged to detect scintillation events emanating from thescintillators.
 11. The radio frequency coil as set forth in claim 10,wherein the scintillators are formed as pixels sized to align with gapsof a conductive screening mesh defining a radio frequency shield orscreen of the resonant structure.
 12. The radio frequency coil as setforth in claim 10, wherein the resonant structure includes conductorsdisposed on a printed circuit board, and the scintillators are disposedon or in the printed circuit board.